The present invention relates to apparatus and methods of reducing scatter such as Compton scatter in radiation imaging systems such as gamma cameras and positron annihilation cameras. More particularly, the present invention relates to various apparatus and methods some of which reduce Compton scatter by subdividing the photopeak window into lower and upper portions and subtracting the lower portion from the upper portion for each of many regions of an image.
Radiation imaging apparatus such as scintillation cameras are well known devices used in nuclear medicine. A radiopharmaceutical (e.g. technetium 99m (Tc-99m), thallium 201 (Th-201), gallium 67 (Ga-67) or indium 111 (In-111)) is administered to the patient and temporarily accumulates in areas of the body where it is desired to image lesions. The radioactive isotope emits gamma rays (or positrons which decay into oppositely directed pairs of gamma rays) which then impinge on a scintillation detector such as sodium iodide adjacent to which is located one or more photomultiplier tubes or other means of converting scintillations to electrical pulses. Electronic apparatus is used to process the pulses and to determine coordinate position values of the source of each gamma ray which are used in forming an image.
The radioisotope source in other applications is located outside the patient and produces a fan beam or a cone beam that passes through the patient or specimen and impinges on an extended detector as in bone mineral densitometry for osteoporosis.
The prior art has recognized that gamma rays produced by radioactive decay may travel directly to the detector or may be scattered by mechanisms such as Compton scattering. When a gamma ray initially emitted in a particular direction is scattered, a gamma ray of somewhat different energy travels in another direction which then reaches the detector as information which interferes with or can be confused with gamma rays reaching the detector directly. Therefore, if the camera is designed to determine that the source of a particular gamma ray lies along a line coinciding with its direction of incidence on the detector, then Compton scatter photons will be misinterpreted as indicating a source distribution of the radionuclide in the patient which is actually different than its actual distribution. In practice the effect of Compton scattering is to provide fuzzy and indistinct images instead of desirable distinct images of the radionuclide distribution in the patient. For example, sodium iodide as a camera detector has a relatively poor energy resolution, and images obtained with conventional techniques employing a pulse height analyzer (PHA) window centered around the photopeak portion of the radiotracer spectrum will contain high (20-50%) fractions of scattered photon events.
As a result, devices designed to image the photon emissions from these radioactive materials are presented with a geometrically widespread mixture of scattered photons and non-scattered photons. There is no geometric method to separate scattered non-scattered photons since the usual source of radioactivity is distributed over a significant volume within the patient, and the emission of photons from radioactive decay is isotropic. The inclusion of scatter photons in the measurements obtained with imaging devices not only degrades the image contrast, but also the degradation is amplified by image reconstruction in emission computed tomographic imaging.
The physics of Compton scattering is well known and the energy of a scattered photon is given by the equation ##EQU1## where e.sub.s is the energy of the Compton scattered photon e.sub.o is the energy of the original emitted gamma ray, mc.sup.2 is the equivalent energy of a particle such as an electron which scatters the original gamma ray (mc.sup.2 for an electron is 511 KeV), and A is the Compton angle or angle between the direction of incidence of the original photon and the direction of travel of the scattered photon.
A scintillation camera typically has one or more pulse height analyzers (PHAs) by means of which the energy of radiation can be readily determined. Gating circuits detect incident photons upon the detector which have radiation energies within a predetermined band or "window" and all of the other radiation not within the window is excluded for imaging purposes. By inspection of the Compton scattering formula (1) it is apparent that a window which is set for a particular energy also determines the Compton angle of any scattered radiation that may enter. For example, a given radionuclide produces gamma rays having a particular energy e.sub.o. Due to statistical factors the actual energy which is sensed at the detector will lie in a more or less peaked distribution centered on an energy called the photopeak energy and the distribution is conventionally called the photopeak. In a typical gamma camera a photopeak window or energy window is centered so as to have lower and upper energy boundaries bracketing the center of the photopeak e.sub.o.
Unavoidably some Compton scattering produces radiation that is within the photopeak window which therefore is confused by the radiation imaging apparatus with gamma rays which have come directly from the radionuclide distribution in the patient. The boundaries of the photopeak window define the maximum angle of Compton scattering which can occur and still lie within the photopeak window and thus be confused with the true source distribution. The Compton angle is determined by solving the Compton formula (1) for Compton angle A after substituting the photopeak energy e.sub.0 and the lower boundary of the photopeak window for e.sub.s.
In one prior art approach a second entirely separate PHA window or energy range distinctly below the photopeak window is provided at lower energies where the Compton scattering is more numerous to acquire a scatter image. An empirically determined proportion of Compton events is subtracted from the number of detected incident gamma rays in the photopeak window as a rough means of estimating the actual Compton scattering in the photopeak and thus reducing its blurring effect on the image. In emission computed tomography (ECT), the scatter image may be used to construct a scatter ECT image which is subtracted from the regular ECT image at the same geometric position.
Unfortunately, the amount of Compton scattered radiation in a lower energy window in general bears little relationship to the actual amount of Compton scatter in the photopeak region itself. Compton scattering occurs due to radiation that originates in regions of the specimen adjacent to the region which is desired to be imaged and the distribution of the radionuclide in any given specimen is precisely the unknown which is to be imaged. As put by one worker
"the contribution of adjacent regions over a particular site to the measured activity is almost entirely in the scatter region when collimated detector systems are used. Thus, the peak-to-scatter ratio would be diminished by the presence of appreciable amounts of radioactivity in adjacent structures. This localization must be determined in the course of mapping the peak-to-scatter ratio in and around the region(s) of interest, and the scatter contribution from the adjacent region(s) taken into account . . . [I]t appears that a more comprehensive formalism is needed, since an almost infinite number of phantom configurations and calibrational studies would be needed . . . PA1 The suggestion that the scatter window be used as an index of the scatter contribution contained in the photopeak region seems destined to failure because of the phenomena described above. If all the scatter could be related to the activity distribution along the axis of the collimator, the procedure . . . might be useful. Since, however, a variable but large fraction of scatter counts arise from surrounding regions, a single correction factor would not be very useful." R. E. Johnston et al., "Inherent Problems in the Quantitation of Isotope Scan Data" Medical Radioisotope Scintigraphy, Vol. I, IAEA, Vienna, 1969, pp. 617-631, at pp. 630-631.
A Compton scattering energy window approach is also discussed in "Four-View Computer Scintiscanning: Image Structuring Through Multi-Window Pulse-Height Analysis" by S. Genna et al. in Medical Radioisotope Scintigraphy 1972, IAEA, Vienna 1973, pps. 133-154 and R. J. Jaszczak "Scatter Compensation Techniques for SPECT", IEEE Trans. Nuclear Science, Vol. 32, February 1985.
In another approach to reducing Compton scatter a gamma camera has scintillation events displayed at X,Y coordinates on a cathode ray tube by unblanking the tube with z pulses applied to its control electrode. A Compton scattered radiation deemphasizer determines where the peaks of total energy pulses fall in a part of the energy spectrum or window which is the photopeak window for present purposes. The deemphasizer causes small Z pulses to be produced in the part of the spectrum where Compton scatter is most prevalant and causes increasingly larger Z pulses as the total energy increases up to the midpoint of the photopeak window where Compton scatter is less significant. Constant amplitude Z pulses are produced at and after the midpoint of the photopeak window. In FIG. 3 of this patent a distribution of energy at the photopeak window and the method of deemphasis is shown. A microprocessor implementation for accomplishing the deemphasis is shown in FIG. 6 of U.S. Pat. No. 4,258,428 to E. M. Woronowicz.
In U.S. Pat. No. 4,575,810 a "simplistic" and "completely synthetic" example (col. 4 lines 29-47) features three windows. The patent asks the reader to suppose that events in an upper section of the photopeak were known to provide twice as much image information as those in the lower section, and events in a region just below the photopeak could be subtracted to remove scatter. An upper window image would be added twice and the lower window image added once, while subtracting the below range window image. However, the patent states that it is likely that signal-to-noise ratios in the weighted image would not be larger than normal images, due to the simplistic nature of the example, and instead teaches an approach using a carefully constructed weighting function.
In U.S. Pat. No. 4,415,982 to M. Nishikawa a background suppression approach to improving the images is produced and no suggestion that scatter reduction is specifically intended is present. In the Nishikawa patent a scintillation camera has a memory with addresses corresponding to the elements of a matrix-like divided image of a scintigram and these addresses correspond to incident position signals of points of radiation determined by a position calculating circuit. A pulse height analyzer receives the radiation events registered by the camera head and issues an unblanking signal to control apparatus if the count rate is significantly higher than background noise. The content at the designated address in memory is compared with a predetermined minimum value of radiation in a comparator. When the content is less than the minimum value the control apparatus adds one to the content and restores the increased content to the designated address. When the content is at least equal to the minimum value, the control apparatus issues the unblanking signal to display apparatus to display the radiation at the designated address. In effect this apparatus accomplishes a kind of subtraction so that only those image positions which receive a number of events in excess of a predetermined number are displayed at all on the display apparatus.
The various approaches of the prior art evidently provide some reduction of the Compton scatter blurring of images and increase the contrast and distinctness of the images somewhat. However, the subtraction approach using a separate Compton window has been criticized in the cited literature because such approach is dependent on particular source distributions and is dismissed in the cited U.S. Pat. No. 4,575,810 as having low signal-to-noise ratio. A deemphasizer as described in U.S. Pat. No. 4,258,428 allows some Compton scatter to be detected because it passes all z pulses below the midpoint of the photopeak window except that these are deemphasized. The scintillation camera of Nishikawa in U.S. Pat. No. 4,415,982 does not purport to reduce Compton scattering. Compton scatter which is included in regions of the image which exceed the predetermined minimum of radiation which is blanked out by Nishikawa will not be reduced at all.